Early and accurate diagnostics are crucial to the treatment of many medical disorders. High-throughput biosensing is an active and growing area of research for clinical applications in detecting antibodies, genes, drugs, peptides, cells, and other biological molecules of interest, as well as for sequencing DNA molecules.
Currently, the two major optical approaches to biosensing upon which research efforts are focused are the use of segmented evanescent waveguide biosensors and the use of the so-called “biochips”, which basically include microfabricated substrates with capture biomolecules secured thereto.
Optical waveguides are often used in low sample concentration, high throughput immunoassays (IA) and molecular diagnostic assays (MDx). When a segmented waveguide is used, the waveguide is illuminated with one or more wavelengths of electromagnetic radiation (e.g., light) to facilitate a determination of presence, absence, or amount of one or more particular analytes of interest in a sample. Numerous optical techniques have been developed to employ the evanescent wave from optical waveguides in biosensing applications.
For example, the use of optical waveguides for high-throughput sensing can be achieved using either mass (Silzel et al., 1998) or fluorescence (Stimpson et al., 1995; Watkins et al., 1998; Plowman et al., 1999) sensing techniques. Both of these sensing techniques have been demonstrated as useful in waveguides for conducting immunoassays (Silzel et al., 1998; Wadkins et al., 1998; Plowman et al., 1999) and genetic screenings (Stimpson et al., 1995).
When an optical waveguide is used as a so-called “mass sensor”, the presence of a captured analyte may be detected or measured by, first, measuring a baseline absorption or refractive index of a capture molecule on or adjacent a surface of the waveguide prior to exposing the capture molecule to a sample and, second, following exposure of the capture molecule to a sample, determining the difference in absorption or refractive index of the capture molecule and any analyte bound, or hybridized, thereto. With this mechanism, the measured signal changes upon hybridization of capture molecules with analytes in the sample. The signal change is in proportion to the mass that lies within the evanescent field of the waveguide.
Waveguide mass sensing techniques typically rely on surface plasmon resonance (SPR) 5, which has been used extensively in optical biosensing (Liedberg et al., 1987). A surface plasmon can exist at the interface between two media, one of which has a negative dielectric constant (Peyghambarian et al., 1993), such as a metal, and can be resonantly excited using a ruled optical grating or prism to obtain phase matching. The index perturbation of the analyte disturbs this resonance, with greater perturbation provided with larger molecular weight. These sensors have been used for immunoassay (Cullen et al., 1987; Morgan and Taylor, 1992) and molecular diagnostic assay (Watts et al., 1995; Nilsson et al., 1997; Bianchi et al., 1997), and are commercially available (see Malmqvist and Karlsson, 1992, for a review). Affinity sensitivities can be in the nM range (Morgan and Taylor, 1992) for large analyte molecular weight, and 50 μM (Karlsson and Stahlberg, 1995) for small analyte molecular weight, and a sensitivity per unit area of 20 fM/mm2 was measured using resonant mirrors (Watts et al., 1995). Nanoparticles coated with the analyte have also been used to increase mass, leading to a sensitivity of 0.1 pM (Kubitschko et al., 1997). Other waveguide mass sensors are based on light scatter (Stimpson et al., 1995), interferometry (Schneider et al., 1997), and waveguide absorption spectroscopy (Mendes and Saavedra, 1999).
When fluorescence sensing is employed, electromagnetic radiation may be used to create an evanescent wave within an optical waveguide that excites a fluorescent dye, which is also referred to as a fluorophore tag, or a similar tag bound, for example, to a molecule that competes with an analyte of interest for a binding site on a capture molecule immobilized on or adjacent to a surface of the waveguide. The fluorophore tag gives off emitted electromagnetic radiation, the intensity of which is indicative of the presence, absence, or amount of the analyte in the sample.
For optical waveguide sensors that employ fluorescence sensing techniques, the sensitivity depends on affinity strength between each analyte and its corresponding capture molecule, as well as upon the absorption coefficient and fluorescence quantum yield of the fluorophore tag. Fluorescence sensing techniques are generally more sensitive and more specific than mass sensing techniques.
Many of the approaches to fluorescence sensing techniques in optical waveguides are based on the use of optical fiber (Abel et al., 1996; Squillante, 1998), but work has also been performed with planar waveguides (Zhou et al., 1991; Herron et al., 1993; Plowman et al., 1996). Planar waveguide sensors have been used in immunoassay (Zhou et al., 1991; Herron et al., 1993) and molecular diagnostic assay (Plowman et al., 1996) studies, where the latter demonstrated low fM sensitivity.
Nanoparticles have been used as light scattering elements (Yguerabide and Yguerabide, 1991) or as an alternative to fluorophore tags.
When optical waveguides are used, the waveguide may include independent sensing zones on or adjacent to which different types of capture molecules are immobilized. All of the discrete sensing zones of a segmented waveguide can be interrogated in parallel by use of a charge-coupled device (“CCD”) array, which can capture the full time dynamics of the affinity interaction. The array size is limited by the patterning of immobilized capture molecules (Silzel et al., 1998) and the sensitivity is limited by the number of captured analytes per sensitivity zone and by the sensing technique employed. Nonetheless, with new patterning techniques, the densities of sensing zones of optical waveguides are continuing to increase (Morgan et al., 1995; Stimpson et al., 1998).
In the state of the art, the sizes of the sensing zones of segmented waveguides are ever-decreasing. Accompanying the decrease in sizes of the sensing zones of segmented waveguides are proportional reductions in the sensitivities with which analytes can be detected by the waveguide. Accordingly, the system requirements of segmented waveguides are becoming ever more stringent.
Moreover, the use of a segmented waveguide is somewhat undesirable since the excitation radiation is not confined within the plane of the waveguide. Consequently, the electromagnetic radiation emitted from one sensing zone may interfere with the electromagnetic radiation emitted from one or more adjacent sensing zones, thereby reducing the optical efficiency of a segmented waveguide and, thus, the degree of confidence with which each analyte may be detected when the segmented waveguide is used. In order to reduce the interference between adjacent sensing zones of segmented waveguides to acceptable levels, the area of each sensing zone to which capture molecules are secured or to which a sample is introduced may be limited to less than 25% of the total area of the sensing zone. As a result, segmented waveguides are relatively insensitive.
A related approach to the optical waveguide is the so-called “biochip”, of which the so-called “DNA chip” (Vo-Dinh et al., 1999; Hacia, 1999; Lipshutz et al., 1999) is an example. Biochips, which may be fabricated using self-assembled monolayer or similar patterning techniques, may have very large arrays of sensing zones (e.g., up to 100,000 or more sensing zones). A sample is applied to a biochip, one or more analytes of interest in the sample bind capture molecules on the biochip, and the presence, absence, or amount of each analyte or the hybridization characteristics of the capture substrate with a corresponding analyte is detected in much the same manner as that which is used when segmented waveguides are employed to detect analytes.
For example, fiber optic probes, scanning near-field microscopes, or confocal microscopes may be used to direct one or more wavelengths of electromagnetic radiation into each sensing zone of a biochip to excite fluorescent dyes within the sensing zone. The fiber optic probe or confocal microscope may then be used to detect the electromagnetic radiation emitted from fluorescent labels within each sensing zone. If a fiber optic prove or confocal microscope is used, the biochip may be raster-scanned, one sensing zone at a time. The sensing zones of a biochip can be closely spaced since the probe itself provides lateral optical confinement, which leads to the possibility of biochips with very high sensing zone densities. Such sequential detection is somewhat undesirable, however, because of the time required to scan a biochip with a dense array of sensing zones. Moreover, each hybridization reaction occurring on a biochip is typically analyzed when it reaches an endpoint, which may take hours. Further, different reactions may require different hybridization temperatures. As a result, it may be difficult to simultaneously effect a number of different hybridization reactions on a single biochip, which further increases the amount of time required to obtain results from raster-scanned biochips.
While parallel detection techniques may also be used to simultaneously analyze multiple sensing zones on biochips, parallel detection techniques may also be complicated by the different temperature dependences of different hybridization reactions to be conducted on a single biochip (Fotin et al., 1998). Again, it may be necessary to carry out reactions and to take measurements at a number of temperatures.
In response to these temperature dependence problems, assay techniques have been developed wherein the hybridization rates of different hybridization reactions occurring at different sensing zones of a biochip are simultaneously measured at one or more distinct temperatures or narrow temperature ranges (Jensen et al., 1997).
Another assay technique has been proposed includes effecting different chemical reactions on the surfaces of microspheres (Micheal et al., 1998). The microspheres are deposited into individual wells etched into the distal end of an imaging fiber bundle. Cavity effects of the wells are masked by incoherent illumination, but the use of a CCD to sense reactions on the surfaces of the microspheres may be used in parallel with the imaging fiber bundle.
Flat, or planar, cylindrical microcavities have been used in low-threshold lasers (McCall et al., 1992; Zhang et al., 1995; Baba, 1997) and optical spectral filters (Rafizadeh et al., 1997; Blom et al., 1997; Madsen and Jhao, 1998; Little et al., 1998). These cylindrical microcavities have been fabricated from a variety of materials, including semiconductor materials (e.g., silicon) and glass. Studies have shown that planar cylindrical microcavities with cavity diameters of 10.5 μm may have free spectral ranges (FSRs) of greater than 35 nm and cavity Q values of greater than 8000. Finite-difference time-domain (FDTD) studies (Li and Liu, 1996; Hagness et al., 1997) performed with such planar cylindrical microcavities suggest that FSRs exceeding 100 nm and cavity Q values of 104 or greater are possible with optimized designs. A theoretical study (Little and Chu, 1996) of cavity surface roughness supports these conclusions. Nonetheless, cylindrical microcavities have typically been illuminated from the planar ends thereof, which causes the illuminating electromagnetic radiation to travel throughout the volume of these cylindrical microcavities.
The inventors are not, however, aware of the use of biosensors that include resonant optical cavities that facilitate the use of whispering gallery modes to provide enhanced sensitivity and that have quality factors of at least about 104.